Ultrasound imaging catheters typically produce cross-sectional views of internal features of a living body to assist in accurate medical diagnosis and treatment decisions. Intravascular and intraluminal ultrasound imaging refers generally to imaging from within blood vessels, arteries, and other small lumens. Intracardiac imaging refers to imaging the walls of the heart. Ultrasound imaging catheters generally image features a short distance from the transducer.
Ultrasound imaging devices are generally known, and although there are various ways to accomplish this type of imaging, imaging catheters employ mechanisms to transmit scanning beams of ultrasound energy into the area being studied and to receive the return echoes from each scan in order to generate an image which can be seen on a visual display such as a monitor. The transmitted ultrasound beam has several important parameters, such as frequency, beam-width in the near-field and far-field, focal length, transition length between near and far-field, and beam intensity which define the performance characteristics of the ultrasound imaging system. Some of the most important performance characteristics of an ultrasound imaging system include axial and lateral resolution and penetration depth. The resolution of an ultrasound imaging system is defined as the minimum distance of separation between two objects and the maximum distance from the two objects at which the two objects can still be identified separately in the ultrasound image. The penetration depth is the depth below the surface of the feature at which the ultrasound waves can detect and produce an image.
Axial resolution and lateral resolution typically improve with increasing frequency. Increasing the frequency of a system, however, reduces the penetration depth of the ultrasound beam and consequently the depth of view. Hence, the optimal acoustic frequency for any given imaging application using conventional ultrasound transducers is a compromise between resolution and penetration depth. Reducing the beam diameter also improves lateral resolution. For an exhaustive discussion of ultrasound beam characteristics and the relationship of these characteristics to performance parameters, the reader is referred to Harm ten Hoff, "Scanning Mechanisms for Intravascular Ultrasound Imaging; A Flexible Approach"; ISBN:90-9006072-3; Ph.D. Thesis (1993); Erasmus University Rotterdam, The Netherlands; "Diagnostic Ultrasound; Principles and Instruments," 4th Ed., Frederick W. Kremkau, W. V. Saunders Co., Philadelphia, ISBN:0-7216-4308-6 (1993); these and all other references cited herein are expressly incorporated by reference as if fully set forth in their entirety herein.
The penetration depth of an ultrasound imaging system is a function of the frequency, the dynamic range of the system, and the ultrasound beam shape. Usually, however, the dynamic range is limited because excessive intensity can result in cavitation which can cause severe damage to the catheter and to the surrounding tissue. The penetration depth can be improved by optimizing the beam shape by extending the beam focal length and producing a tighter beam at a desired distance from the transducer. Hence, imaging system design aspects other than frequency may be changed to affect the beam shape thereby improving the resolution and penetration depth of the ultrasound imaging system.
In turn, the shape of an ultrasound beam of a single transducer is a function of the shape and dimensions of the transducer, the ultrasound frequency and the use of focusing. Referring to FIG. 1, a typical ultrasound beam-shape is described. The near-field extends from the transducer and is characterized by a converging beam and an irregular ultrasound intensity pattern due to interference of ultrasound waves originating from different parts of the transducer. In the near-field, the acoustic field amplitude and phase are erratic due to the constructive and destructive interference of energy emitted by different parts of the transducer. The near-field can be described as somewhat incoherent.
The transition from near-field to far-field conditions is located at a distance Z=Z.sub.R from the transducer. The transition length, Z.sub.R, is defined as the distance from the transducer along the central axis of the transducer where the difference between the acoustic path-length from the center of the transducer and from the periphery of the transducer is one-half a wavelength (1/2.lambda.). This can be better described with reference to FIG. 11 which depicts a transducer having a flat acoustic element 2 and no focusing lens. Throughout this application, "standard transducer" refers to a transducer having a flat acoustic element and no focusing lens. In FIG. 11, the lines b and Z.sub.R represent acoustic path-lengths. The acoustic path-length is measured from an iso-phasic (coherent) plane on or inside the transducer assembly, usually the front surface of the acoustic element. The acoustic path-length through a media can be expressed in the number of wavelengths and is equal to: ##EQU1##
where L.sub.x =geometric path-length through media.sub.x
c.sub.x =speed of sound in media.sub.x PA1 f=acoustic frequency
The distance b represents the acoustic path-length from the periphery of the transducer, and the distance Z.sub.R represents the acoustic path-length from the center of the transducer. The transition length, Z.sub.R, is located at a distance from the transducer where: EQU b-Z.sub.R =1/2.lambda.
Using this definition, it can be seen that closer to the transducer, in the near-field, the difference, b-Z.sub.R, is larger than 1/2.lambda.. Conversely, further away from the transducer, in the far-field, the difference, b-Z.sub.R, is smaller than 1/2.lambda., resulting in an increasingly coherent and diverging beam.
The far-field is distinguished by wavefronts (iso-phasic lines) which are convex-shaped and diverging. The natural focal point is located at the point of minimum beam width, W.sub.f, which is at a distance from the transducer defined as the focal length, Z.sub.f, generally about 0.8Z.sub.R from the transducer surface.
According to the above definitions, Z.sub.R and the natural focal length Z.sub.f will increase when the acoustic path-length from the periphery of the transducer increases with respect to the acoustic path-length from the center of the transducer. Hence, increasing the diameter of the transducer will increase Z.sub.R and also Z.sub.f. By the same principle, Z.sub.R and Z.sub.f can be decreased by decreasing the acoustic path-length from the periphery of the transducer with respect to the acoustic path-length from the center of the transducer.
An example of the effect of shortening Z.sub.R and Z.sub.f is seen in the known use of a concave lens to focus an ultrasound beam. Referring to FIG. 2, an ultrasound transducer comprises a flat acoustic element 2 having a concave lens 3 laid on the front of the acoustic element. Z'.sub.R and Z'.sub.f represent the transition length and focal length, respectively, of a standard transducer. The material of the lens is selected so that the speed of sound in the lens, c.sub.1, is greater than the speed of sound in the medium (usually water/blood), c.sub.m. Because the path, b, from the periphery travels a greater distance through the lens at a greater speed than the path, a.sub.R, from the center, the acoustic path-length from the periphery of the transducer is decreased relative to the acoustic path-length from the center of the transducer to the transition point, Z'.sub.R, of a same sized standard transducer. Accordingly, the transition length Z.sub.R and the focal length Z.sub.f are decreased relative to Z'.sub.R and Z'.sub.f for the transducer without the lens.
A similar effect can be observed when the acoustic path-length along the axis is shortened with respect to the acoustic path-length from the periphery. In other words, when a portion of the acoustic path-length along the axis is reduced, the remainder of the acoustic path-length must be increased, thereby increasing the transition length Z.sub.R to maintain the 1/2.lambda. difference between the acoustic path-lengths from the periphery and from the center of the transducer.
As the beam travels through the near-field, its diameter decreases, and the beam is described as being tight. As the beam travels through the far-field, its diameter increases and the beam is described as diverging. This is true even for flat unfocused transducer elements because even for flat transducer elements there is some beam narrowing or "focusing." For a given transducer diameter, increasing the beam frequency increases the near-field length and produces a narrower beam. Reducing the beam diameter improves the lateral resolution. Therefore, lateral resolution is greatest at the focal point and decreases beyond the focal point as the beam travels through the far-field in a diverging pattern.
A larger transducer may also improve lateral resolution and penetration depth, but a larger transducer requires a larger catheter, and the catheter size may often be restricted by the size of the vessels through which it must fit.
As described above, with reference to FIG. 2, an ultrasound beam shape may also be adjusted by the use of focusing. Ultrasound waves have been disclosed to be focused by using a curved (rather than flat) transducer element, a curved reflector in the transducer assembly, or a lens. The emphasis of ultrasound imaging design has been concentrated on the ability to image smaller and smaller vessels in the distal regions of the coronary vasculature via a remote arterial site to obtain images of the vessel walls. This has been accomplished by inserting a catheter having an ultrasound transducer disposed on its distal end into a vessel and positioning the transducer proximate to the area of the vessel wall to be imaged. To be inserted into very small vessels, the size of the ultrasound transducer must be relatively small not only to traverse the vessels, but also to avoid occluding the vessel. Because the transducer is placed immediately adjacent or very close to the area to be imaged, a short focal length is required.
To this end, focusing has been used to converge the ultrasound beam for the purpose of shortening the focal length. For example, the ultrasonic transducer described in U.S. Pat. No. 5,438,999 to Kikuchi et al. includes a piezoelectric element having a concave surface to converge the ultrasonic beam, producing a shortened near-field and focal length and a tight beam in close to the transducer. A concave focusing device is used to shorten the focal length of the transducer in order to increase resolution and detection depth in close to the transducer. Kikuchi et al. also discloses an ultrasonic transducer having a convex acoustic lens made of silicone rubber materials for converging an ultrasonic beam. A convex lens made of silicone rubber materials will create a converging beam having a shortened focal length because the speed of sound in silicone (about 1 mm/.mu.sec as shown in FIG. 19e tabulating materials and respective densities and speeds of sound) is slower than the speed of sound in the typical medium of water/blood (1.5 mm/.mu.sec). Hence, the silicone rubber convex lens disclosed in Kikuchi et al. acts to increase the acoustic path-length along the center axis of the transducer with respect to the acoustic path-length from the periphery of the transducer, thereby decreasing the transition length, Z.sub.R, and the focal length, Z.sub.f.
However, the short focal lengths of the conventional flat and concave-shaped transducers limit the imaging distance from the transducer at which features may be imaged with good resolution and detection depth. These limitations restrict the type and size of internal features within the human body which may be imaged by systems utilizing these transducers. For instance, an ultrasound imaging catheter with a small enough diameter to insert intraluminally into the chambers of the heart may not have a large enough focal length to image the internal walls of the chambers with adequate resolution and detection depth.
More specifically, the maximum size catheter which can be inserted through the femoral artery to the left side of the heart is about 8F, and through the inferior vena cava to the right side of the heart is about 9F or 10F. A 9F catheter can accommodate a transducer up to a diameter of about 1.93 mm. Within the chambers of the heart, it is desirous to image a distance of up to 10 cm in order to image the walls of the heart. At a frequency of about 9 megahertz (MHz), which is needed for adequate penetration, the high resolution near-field for a flat transducer extends to only about 4.5 mm. Therefore, when the ultrasound beam reaches the walls of the heart chambers, the beam is well into the far-field where there is reduced resolution and penetration depth.
A need exists for an improved ultrasound transducer with an extended focus at low frequencies which can be used to image large chambers, organs, vessels, or other anatomic structures with high resolution and high penetration depth.